This invention relates to nuclear magnetic resonance (NMR) imaging methods. More specifically, the invention relates to improved NMR methods for rapidly and accurately imaging one or more of the computed spin-lattice (T.sub.1), and spin-density (M.sub.o) parameters.
By way of background, the nuclear magnetic resonance phenomenon occurs in atomic nuclei having an odd number of protons and/or neutrons. Due to the spin of the protons and the neutrons, each such nucleus exhibits a magnetic moment, such that, when a sample composed of such nuclei is placed in a static, homogeneous magnetic field, B.sub.o, a greater number of nuclear magnetic moments align with the field to produce a net macroscopic magnetization M in the direction of the field. Under the influence of the magnetic field B.sub.o, the magnetic moments precess about the axis of the field at a frequency which is dependent on the strength of the applied magnetic field and on the characteristics of the nuclei. The angular precession frequency, .omega., also referred to as the Larmor frequency, is given by the equation .omega.=.gamma.B, in which .gamma. is the gyromagnetic ratio which is constant for each NMR isotope and wherein B is the magnetic field acting upon the nuclear spins. It will be thus apparent that the resonant frequency is dependent on the strength of the magnetic field in which the sample is positioned.
The orientation of magnetization M, normally directed along the magnetic field B.sub.o, may be perturbed by the application of a magnetic field oscillating at the Larmor frequency. Typically, such a magnetic field designated B.sub.1 is applied in a plane orthogonal to the direction of the static magnetic field by means of a radio frequency (RF) pulse through coils connected to a radio-frequency-transmitting apparatus. The effect of field B.sub.1 is to rotate magnetization M about the direction of the B.sub.1 field. This may be best visualized if the motion of magnetization M due to the application of RF pulses is considered in a Cartesian coordinate system which rotates at a frequency substantially equal to the resonant frequency .omega. about the main magnetic field B.sub.o in the same direction in which the magnetization M precesses (i.e., the rotating frame). In this case, the B.sub.o field is typically chosen to be directed in the positive direction of the Z-axis, which, in the rotating frame, is designated Z' to distinguish it from the fixed-coordinate system. Similarly, the X- and Y-axes are designated X' and Y'. Bearing this in mind, the effect of an RF pulse, then, is to rotate magnetization M, for example, from its direction along the positive Z' axis toward the transverse plane defined by the X' and Y' axes. An RF pulse having either sufficient magnitude or duration to rotate magnetization M into the transverse plane (i.e., 90.degree. from the direction of the B.sub.o field) is conveniently referred to as a 90.degree. RF pulse. Similarly, if either the magnitude or the duration of an RF pulse is selected to be twice that of a 90.degree. pulse, magnetization M will change direction from the positive Z' axis to the negative Z' axis. This kind of an RF pulse is referred to as a 180.degree. RF pulse, or for obvious reasons, as an inverting pulse. It should be noted that a 90.degree. or a 180.degree. RF pulse will rotate magnetization M through the corresponding number of degrees from any initial direction of magnetization M. It should be further noted that an NMR signal will only be observed if magnetization M has a net transverse component (perpendicular to B.sub.o) in the transverse plane. Assuming an initial orientation of magnetization M in the direction of the B.sub.o field, a 90.degree. RF pulse produces maximum net transverse magnetization in the transverse plane since all of magnetization M is in that plane, while a 180.degree. RF pulse does not produce any transverse magnetization. 180.degree. RF pulses are frequently utilized to produce NMR spin-echo signals. RF pulses may be selective or nonselective. Selective pulses are typically modulated to have a predetermined frequency content so as to excite nuclear spins situated in preselected regions of the sample having magnetic-field strengths as predicted by the Larmor equation. The selective pulses are applied in the presence of localizing magnetic-field gradients. Nonselective pulses generally affect all of the nuclear spins situated within the field of the RF pulse transmitter coil and are typically applied in the absence of localizing magnetic-field gradients.
There are two exponential time constants associated with longitudinal and transverse magnetizations. The time constants characterize the rate of return to equilibrium of these magnetization components following the application of perturbing RF pulses. The first time constant in referred to as the spin-lattice relaxation time (T.sub.1) and is the constant for the longitudinal magnetization to return to its equilibrium value. Spin-spin relaxation time (T.sub.2) is the constant for the transverse magnetization to return to its equilibrium value in a perfectly homogeneous field B.sub.o. In fields having inhomogeneities, the time constant for transverse magnetization is governed by a constant denoted T.sub.2 *, with T.sub.2 * being less than T.sub.2.
There remains to be considered the use of magnetic-field gradients to encode spatial information (used to reconstruct images, for example) into NMR signals. Typically, three such gradients are necessary: EQU G.sub.x (t)=.differential.B.sub.o /.differential..sub.x, EQU G.sub.y (t)=.differential.B.sub.o /.differential..sub.y,
and EQU G.sub.z (t)=.differential.B.sub.o /.differential..sub.z.
The G.sub.x, G.sub.y, and G.sub.z gradients are constant throughout the imaging slice, but their magnitudes are typically time dependent. The magnetic fields associated with the gradients are denoted, respectively, b.sub.x, b.sub.y, and b.sub.z, wherein EQU b.sub.x =G.sub.x (t)x, EQU b.sub.y =G.sub.y (t)y, EQU b.sub.z =G.sub.z (t)z,
within the volume.
In the application of NMR to medical diagnostic imaging of a live human subject, for example, the intensity of each picture element (pixel) of the image is a complex function of the tissue-related NMR parameters of the NMR pulse sequence utilized to gather the imaging information. The tissue-related NMR parameters are the afore-described T.sub.1 and T.sub.2 relaxation times, as well as the spin density (M.sub.o) of the particular nuclear species (H.sup.1, P.sup.31, C.sup.13, etc.) imaged. The proton (H.sup.1) is a typically selected nuclear species for NMR study, due to its abundance in water associated with tissues. The pulse sequence timing parameters of importance for image tissue contrast and intensity are the sequence-repetition time T.sub.r and the NMR spin-echo delay time T.sub.e, both of which will be described hereinafter in greater detail.
It is desirable to produce an image in which the intensity of each pixel depends only on T.sub.1 or T.sub.2 or M.sub.o, since such an image would depend only on tissue-related parameters and magnetic-field strength. Hence, such images should be easier to interpret, medically, since the complicated dependence of pixel intensity on, for example, T.sub.r, T.sub.e, and combinations of T.sub.1, T.sub.2, and M.sub.o is reduced to a single, linear dependence. In the past, computed T.sub.1 images have been produced utilizing an approximation to calculate the T.sub.1 values. Due to the approximate nature of the equation utilized in computing T.sub.1, it is necessary to use sequence repetition times T.sub.r which are much greater than spin-echo times T.sub.e to preserve accuracy of the computed values. However, scan times are proportional to T.sub.r, so that small values of T.sub.r are desirable to keep data collection time short. Typical times to acquire image data in NMR are on the order of several minutes so that physiological processes, such as breathing, cardiac motion, or peristaltic motion can cause motion of tissues between the separate application of the pulse sequences utilized to gather the data, hence, making the calculation inaccurate and lead to motion artifacts in the reconstructed images. Additionally, the computed pixel intensity values are modulated by noise and systemic errors present in any actual NMR system. The values computed for T.sub.1, for example, will be in error by an amount related to such noise. A conventional method to improve accuracy is to increase the number of scans to more than two using different values of T.sub.r for each new scan. However, for the reasons stated, the scan times become unacceptably long. It is, therefore, a principal object of the present invention to provide NMR pulse sequences which enable NMR imaging data to be collected rapidly and for constructing computed images having improved accuracy.